FIELD OF THE INVENTION
The present invention relates to cardiac event detection, and more particularly, to circuits and methods for accurately generating a programmable reference signal(s) that may be used by threshold detector circuits within a sensing channel of an implantable cardiac pacemaker.
The major pumping chambers in the human heart are the left and right ventricles. The simultaneous physical contraction of the myocardial tissue in these chambers expels blood into the aorta and the pulmonary artery. Blood enters the ventricles from smaller antechambers called the left and right atria which contract about 100 milliseconds (ms) before the ventricles. The physical contractions of the muscle tissue result from the depolarization of such tissue, which depolarization is induced by a wave of spontaneous electrical excitation which begins in the sinus node of the right atrium, spreads to the left atrium and then enters what is known as the AV node which delays its passage to the ventricles via the so-called bundle of His. The frequency of the waves of excitation is normally regulated metabolically by the sinus node. The atrial rate is thus referred to as the sinus rate or sinus rhythm of the heart.
Electrical signals corresponding to the depolarization of the myocardial muscle tissue appear in the patient's electrocardiogram. A brief low amplitude signal, known as the P-wave, accompanies atrial depolarization. A much larger amplitude signal, known as the QRS waveform complex, having a predominant R-wave, signifies ventricular depolarization. Repolarization prior to the next contraction is marked by a broad waveform in the electrocardiogram known as the T-wave.
A typical implanted cardiac pacer (or pacemaker) operates by supplying missing stimulation pulses through an electrode on a pacing lead in contact with the atrial or ventricular muscle tissue. The electrical stimulus independently initiates depolarization of the myocardial (atrial or ventricular) tissue, resulting in the desired contraction. Advantageously, electrical signals corresponding to the P-wave or R-wave can be sensed through the same lead, i.e., the pacing lead, and may thereafter be used as a timing signal to synchronize or inhibit stimulation pulses in relation to spontaneous (natural or intrinsic) cardiac activity. Such sensed signals are referred to as an atrial electrogram signal (corresponding to the P-wave) or a ventricular electrogram signal (corresponding to an R-wave).
Every modern-day implantable pacemaker includes a sensing or threshold circuit, whether the activity of one or both chambers of the heart are sensed. When the electrical signal of the atrial or ventricular electrogram is coupled from the heart into the pacemaker, the electrical signal is typically passed through a series or group of electrical circuits referred to as the sensing channel. One of the primary functions of the sensing channel is to detect the presence of particular waveforms within the electrogram signal, e.g., the QRS waveform complex, and to generate a trigger signal whenever such particular waveforms are present, and to not generate such trigger signal when such waveforms are not present. In this way, the sensing channel thus monitors the heart for the occurrence of particular cardiac events, e.g., atrial and/or ventricular depolarization.
Heretofore, the sensing channel has typically included the following electrical circuits: a programmable gain filter, a gain decoder, a window comparator and a constant reference voltage generator. Such circuits, in combination, are frequently referred to as the "sense amplifier." The programmable gain filter receives the electrogram signal and, in operation, filters and amplifies the electrogram signal by a prescribed amount before presenting a filtered, amplified electrogram signal to the window comparator. The gain of the programmable gain filter is controlled so that the amplitude of the filtered, amplified electrogram signal will be sufficient to trigger the window comparator whenever the particular waveform to be detected is present within the electrogram signal, and will be insufficient to trigger the window comparator whenever the particular waveform is not present within the electrogram signal. A threshold or window voltage, typically generated by the constant reference generator, is used to define when the filtered, amplified electrogram signal amplitude is of sufficient amplitude to trigger the window comparator. The presence or absence of the trigger signal generated by the window comparator is then used to control the other circuits within the cardiac pacer.
The gain decoder, coupled to the programmable gain filter, generates a gain signal that controls the gain of the programmable gain filter. The gain signal is set (generated) in response to a gain code received from the other circuits of the cardiac pacer. In order to produce the specified gain, the programmable gain filter includes a bank of switched capacitors that are programmably switched in and out, or substituted, for each other within the programmable gain filter's amplifier circuit. As these capacitors are switched in and out of the programmable gain filter circuit, the gain of the filter circuit changes, as is known in the art.
The internal circuits of an implantable cardiac pacer are typically integrated into a single integrated circuit or "chip." The use of such integrated circuit chip offers the advantage of providing a circuit that is much smaller than similar circuits constructed either partially or completely from discrete components.
A sense amplifier realized using an integrated circuit chip typically consists of a programmable gain filter. The gain of such filter is programmed by switching capacitors in and out of the circuit. The gain of any switched-capacitor filter depends primarily on the ratio of two capacitors (usually the ratio of an input capacitor to a feedback capacitor). The absolute values of capacitors used in integrated circuit chips may vary as much as .+-.20% from chip to chip, wafer to wafer, or run to run. Such variations, however, do not necessarily translate to a gain change of .+-.20%. That is, as explained below, the gain is set by the ratio of capacitors, and a variation of one capacitor, e.g., the input capacitor, is almost always accompanied by a corresponding variation in the other capacitors used on the same chip. Hence, even though the absolute values of capacitors used in integrated circuit chips may vary over a wide range, the ratio of two reasonably sized capacitors, e.g., 1-5 pf, may have a tolerance as low as .+-.0.2%. Hence, the use of a switched-capacitor filter has heretofore provided the basis for an accurate programmable gain within the sense amplifier of a pacemaker.
To illustrate, it is noted that the value of a rectangular integrated circuit capacitor is ##EQU1## where C is the capacitor value, E.sub.ox is the permittivity of silicon dioxide (a constant), t is the oxide thickness between the capacitor plates, W is the width of the capacitor edge, and L is the length of the capacitor edge. In a given integrated circuit chip, the capacitance value can thus vary as a function of variations in the width W, length L, or oxide thickness t. Since these parameters are poorly controlled during integrated circuit fabrication, the absolute value of the capacitor may change as much as .+-.20% from chip to chip, or wafer to wafer, or run to run.
Fortunately, for a parameter such as gain, the value of the parameter is a function of the ratio of two capacitors, e.g., an input capacitor C.sub.1 and a feedback capacitor C.sub.2. That is, the gain, G, of the filter may be expressed as ##EQU2## where k is a constant factor. The variation of gain, .DELTA.G, as a function of variations of capacitors may thus be expressed as ##EQU3## Thus, in a specific integrated circuit, where both C.sub.1 and C.sub.2 will almost always increase or decrease by the same amount (because both will normally experience the same variations in width W, length L, and/or oxide thickness t), any parameter such as gain (G) will theoretically not change at all. (The .+-.0.2% indicated above comes from other second order effects, such as randomness of capacitor edges within the integrated circuit.)
The above discussion and analysis is valid only when the values of the capacitors used in the switched-capacitor filter are reasonably high, e.g., on the order of 1-2 pf. That is, when reasonable value capacitors are used (1-2 pf), the variation of edges are insignificant compared to the size of the total capacitor value. Further, even though there are parasitic (unwanted) capacitors that are always present within an integrated circuit chip (due, e.g., to the conductive traces or layers that interconnect the various circuit elements within the chip), the value of such parasitic capacitors is usually small and therefore negligible (or at least tolerable) compared to the desired capacitance value. Typically, parasitic capacitors are in parallel with the desired capacitors, and they thus tend to increase the overall capacitance value. While it is possible to minimize parasitic capacitance by using intelligent routing of the traces used within the integrated circuit layout, it is impossible to completely eliminate them. Moreover, some parasitic capacitors are hard to predict or model. Thus, heretofore, the only effective way of dealing with parasitic capacitors in switched-capacitor filters, and other integrated circuit chips, has been to design the circuits with capacitor values that are significantly higher than the inherent parasitic capacitance.
Unfortunately, there is a drawback to designing high value capacitors (e.g., capacitors on the order of 1-2 pf or greater) into the programmable filter of a sense amplifier of a cardiac pacemaker when such programmable filter is part of an integrated circuit. When the capacitance value is relatively high, the capacitors (as they are switched into and out of the circuit, and as they perform their desired signal processing functions within the designed circuits) must be charged and discharged by the other pacemaker circuits (usually operational amplifiers). Such charging and discharging translates to a higher current consumption for the operational amplifiers (or other circuits). A high current consumption, of course, is not desirable within an implanted pacemaker, because such current consumption translates to a shorter battery life. Moreover, in order to realize a higher gain, the input capacitor value may have to become very large, which means the capacitor physically occupies a significant amount of space ("real estate") on the integrated circuit chip. In order to design of a programmable gain filter on a chip within an implanted pacemaker that has reduced current consumption, and that occupies an acceptable amount of space, it would be desirable to use capacitors having values on the order of 0.1 to 0.2 pf. Unfortunately, for these values of capacitors, the inherent parasitic capacitance on the chip is significant, and can disadvantageously corrupt the gain value of the programmable gain filter. Thus, there is a need in the art for a programmable gain filter design that allows small capacitance values to be used, thereby reducing current consumption and occupying less chip real estate, while still providing an accurate way to programmably control the gain of the filter.
As an example of the type of gain control that is needed within a sense amplifier (programmable gain filter) of an implantable pacemaker, consider a programmed change in the sensitivity level from 7.75 mV to 8 mV. Such a change may require that the gain of the filter change from 20.65 to 20.0, or about 3%. Such a precise gain change in a switched-capacitor programmable gain filter using small capacitance values (0.1 to 0.5 pf) has heretofore been impossible to realize.
The term "sensitivity" when referring to the sensing channel (or sense amplifier) of a pacemaker reflects the proneness of the sensing channel to generate a trigger signal. A more sensitive sense amplifier generates a trigger signal in response to lower amplitude electrogram signals than will a less sensitive sense amplifier. A related parameter to the sensitivity is the "sensitivity level" of the sense amplifier. The sensitivity level is the unamplified amplitude of the electrogram signal at which the amplified electrogram signal triggers the window comparator. Thus, as the sensitivity level of the sense amplifier decreases (i.e., as the minimum amplitude of electrogram signal needed to generate a trigger signal decreases), the sensitivity of the sense amplifier increases (i.e., the sense amplifier becomes more prone to generate a trigger signal). Unfortunately, because the typical window comparator generates its trigger signal based upon a comparison of the amplified filtered electrogram signal with a constant threshold or window voltage, the sensitivity level cannot be increased or decreased in uniform increments, e.g., of 0.25 mV, unless the steps in the gain of the gain filter are inversely adjusted. That is, if the sensitivity level is to increase (i.e., if the sensitivity is to decrease), the steps in gain, or gain steps, must become smaller; and if the sensitivity level is to decrease, the gain steps must become larger.
Problematically, when small capacitance values are used within a switched-capacitor programmable gain filter, as indicated above, it is not possible to achieve the type of gain precision that may be needed. That is, at high sensitivity levels, the programmable gain filter may not have sufficient precision to assure that the window comparator is triggered or not triggered when appropriate. For example, for an increase in sensitivity level from 0.25 mV to 0.5 mV (a 0.25 mV increase in sensitivity level), and assuming a constant threshold or window voltage of 160 mV above which the amplified filtered electrogram signal must reach before a trigger signal will be generated, the gain must decrease from 640 to 320, a decrease in gain of 320, or 100%. However, for an increase in sensitivity from 7.75 mV to 8.0 mV (also a 0.25 mV increase in the sensitivity level) the gain must decrease from 20.65 to 20.00, a decrease in gain of only 0.65, or roughly 3%. As mentioned above, gain changes on the order of 3% are, at best, difficult to achieve under the best of circumstances, but are essentially impossible to realize when small capacitors (0.1-0.5 pf) are used as part of the integrated circuit design.
Furthermore, as also mentioned above, the gain of the programmable gain filter is typically directly proportional to the capacitance of the capacitors that are substituted into the programmable gain filter circuit. Thus, at low sensitivity levels, and thus high gains, the capacitance needed is quite large relative to the capacitance needed at higher sensitivity levels. Such large capacitances are difficult to fabricate using integrated circuit technology and require large chip areas.
A further disadvantage of the existing programmable gain filter is that, in addition to being imprecise, its gain steps may be polytonic. That is, as the gain code is increased, the gain will generally increase, but because of the poor precision in the switched capacitors, some of the gain steps may be negative, causing an unpredictable decrease in the gain. To further complicate this problem, such "negative gain steps" may not appear uniformly in the fabrication of an integrated circuit chip. Thus, several sense amplifiers fabricated together may exhibit negative gain steps in response to different gain codes. The existence of such polytonic operation thus requires individual circuit testing, advanced programming and/or additional circuitry to correct.
It is therefore evident that improvements are needed in the sense amplifier circuits used within implantable cardiac pacers, and similar implantable medical devises.